Nuclear Magnetic Resonance Imaging (henceforth: MRI), in particular for imaging of selected portions of a patient or another object, has so far relied on two main categories of radiofrequency (RF) coils, as generally described in U.S. Pat. No. 4,920,318. The first type is a volume resonator dimensioned to be disposed around the entire object or patient to be imaged or around a portion thereof. The other type of coil is generally formed by wrapping wire or other conductors on a flat dielectric sheet shaped in such way to be positionable adjacent the portion to be imaged.
There is a strong trend in MRI technology to go to ever higher static magnetic field strengths. The benefits of this trend are a stronger intrinsic signal yield due to the stronger polarization, increased contrast mechanisms such as BOLD or phase contrast, and greater spectral separation for spectroscopic imaging. Since MRI is generally a technique limited by signal to noise ratio (SNR), the increase in SNR and contrast achievable by increasing the static magnetic field strength enables higher spatial and temporal resolution.
However, an increase in static magnetic field strength requires a concomitant increase in operating radiofrequency. As a consequence, the wavelength of the RF radiation is shortened correspondingly. For example, the vacuum wavelength of the RF used on a 7 Tesla system is 1 m, and it can shrink down to 10 cm in human tissue due to the high permittivity of the material. Since the so-called near field of a conductive structure generally scales with the wavelength under consideration, the near field domain of conventional MRI RF probes when used at such high static magnetic fields will shrink below the typical sample size in human MRI. Moreover, as such RF probes are targeted to work with the sample in their near field, they reach their operating limits and have bad prospects for application if one were to go to even higher field strengths.
In practice, the problems arising when the resonant structures presently used for nuclear magnetic resonance (henceforth: NMR) signal probing are taken to high field strengths are a loss of efficiency and—more detrimentally—a spatially inhomogeneous coupling and a reduction of the size of the volume that can be imaged. Most of these problems arise due to interference effects of the RF electromagnetic field emitted and/or received by the probe. For instance, constructive interference yields a signal overshoot in the center region of brain scans obtained using standard resonators, while temporal regions of the head suffer from destructive interferences. In the case of transmission, these field inhomogeneities can induce changes in contrast throughout the image, thus impeding diagnostics. Furthermore, since the sample is placed in the near field domain of the probe, the interaction between probe and sample is strong and increases with higher frequency.
Moreover, the probe design becomes more intricate because resonator and coil design suffers from reduced robustness due to variable loading in in-vivo applications and because current distributions on the conductive structures of the probe are harder to control.
Finally, safety aspects become very intricate due to the strong coupling between sample and probe. RF induced heating is one of the major concerns, especially at higher frequencies. Slight changes of patient geometry next to the probe conductor can crucially change the situation, and worst case scenarios are hard to determine. Safety validation thus becomes a time consuming step in the development of new MRI probes, especially for transmit probes.